The need for sub-surface tissue imaging for medical diagnostics and image guidance and control of therapeutic and surgical procedures is well recognized. Often such imaging has to be performed in a complex, branching network of narrow and difficult-to-reach body lumens (such as, for example, blood vessels of cardiovascular and neurovascular systems, airway tree of lungs, gastrointestinal, bile and urinary tracts) or in tight spaces of natural or surgically created body cavities. There exist endoscopic ultrasound imaging devices, for example intravascular ultrasound (IVUS) or endobronchial ultrasound (EBUS) that address such need.
A higher spatial resolution imaging modality referred to as optical coherence tomography (OCT) has been recently developed and applied for endoscopic imaging in body lumens as well. The OCT systems are numerous and utilize short coherence length light sources for time-domain (TD-OCT) and spectral domain (SD-OCT) incarnations of the OCT, frequency tunable light sources for frequency domain (FD-OCT) version of OCT (see, for example, U.S. Pat. Nos. 5,321,501; 6,134,003).
OCT systems have been described that employ small diameter forward-looking probes with a push-pull actuation scheme (U.S. Pat. Nos. 6,445,939; 7,848,791); and that utilize catheters directed to location of interest thought the use of a guide wire under fluoroscopic guidance with C-arm fluoroscopic equipment (U.S. Pat. Nos. 5,321,501, 6,134,003 and 6,445,939). Fluoroscopic guidance, however, becomes cumbersome and unpractical to use in a branching, three-dimensional (3D) network of body lumens as the C-arm needs to be constantly re-adjusted. Additionally, the fluoroscopic guidance approach lacks sufficient resolution and contrast needed for navigation in small lumens, and exposes patients to harmful ionizing radiation. Even in relatively larger lumens of a cardiovascular system with simple branching, the fluoroscopic guidance is known to have problems with differentiating between main and collateral vessels during placement of catheters, especially in the case of total or partial vessel occlusion.
The use of an endoscopic image guidance and navigation and advancement of imaging probes through the endoscope working channel (described for example in U.S. Pat. Nos. 6,069,698 and 6,564,089) employs typical endoscope system for imaging and a working channel for suction and/or tools the overall diameter of which is about 5 or 6 mm, which limits the practical application of such system to large lumens. Indeed, only about one third or one fourth of bifurcation level of an airway tree in the lungs can be reached with commercially available bronchoscopes.
Imaging probes can be delivered to target locations in reliance on position sensors integrated into distal ends of separate navigating probes (U.S. Pat. No. 7,233,820) or of the imaging probe itself and using virtual images reconstructed from prior-obtained data from computer tomography, CT, or other imaging modalities. However, such a priori CT data is simply unavailable in many clinical situations, for example, in emergency care. Furthermore, the accuracy of so-defined navigation (according to U.S. Pat. No. 7,233,820 itself) depends on accuracy of the position sensors, accuracy of registering the CT data with the tree-dimensional reference of frame, and accuracy of registering the CT data with moving body of patient. These accuracy parameters are typically insufficient for small and peripheral lumens. The additional use of CCD or CMOS image sensors to obtain three camera views from different positions to improve registration of CT data with a 3D reference frame (also disclosed in U.S. Pat. No. 7,233,820) is based on triangulation of two-dimensional (2D) camera images and, therefore, has limited accuracy. In addition, integration of a camera and a position sensor does not allow achieving further miniaturization of the distal portion of the probes.
These deficiencies of fluoroscopy-based guidance and/or guidance relying on auxiliary endoscopes would be at least partially reduced if the imaging probes had their own integrated means for 3D steering and used built-in imaging capabilities for guidance and navigation. These imaging capabilities should include either capability for sufficiently long range (larger field and depth of view) imaging to image sufficient anatomic structure for reliable navigation or capability to constantly register images of local anatomic structure obtained at the distal end with global position referenced to gross anatomy of the imaged organ and/or patient body.
Sub-surface endoscopic imaging modalities referred to in related art possess certain shortcomings including a trade-off between imaging resolution and imaging depth and a trade-off between imaging resolution and probe's insertion widths. Probes having small insertion widths are, additionally, deficient with respect to directions in which such small-width probes are enabled to view the ambient medium. The deficiency stems from common implementation of side-view radial or spiral scanning, which naturally lends itself in enhancing the side looking capability of such probes but not the forward looking capability. The application of the rotational motion to enhance the forward looking capability remains a desirable goal.
One trade-off is imposed by limitations of penetration depth. While ultrasound imaging (US) has resolution on the order of 100 um, which is inferior to that of OCT (about 2 to about 10 um), the penetration depth of the ultrasound is about 10 mm, which is superior to that of about 1 to about 2 mm typical for OCT. The combination of endoscopic OCT and U.S., has been discussed, for example, in U.S. Pat. No. 7,935,060. The combination, in one imaging probe, of optical-image-forming components with ultrasound transducers and wires to produce co-registered OCT and US images increases the size of the probe and effectively prevents further miniaturization of the probe's distal end. In addition, ultrasound imaging cannot be effectively used with imaging through air-gaps, thereby requiring physical contact between probe and the tissue and/or use of liquid-filled balloons. Finally, the need for fluoroscopic and/or endoscopic guidance of a probe serves to disadvantage of the combination probes disclosed in U.S. Pat. No. 7,935,060 and US 2011/0098572.
Another trade-off between the resolution and imaging depth for the probes, that use scanning of focused optical beams, is imposed by limitations of Gaussian optics. Specifically, while the axial (depth) resolution of OCT imaging can be as small as few microns and is determined by properties of light sources, the lateral resolution, i.e., spot size, is typically few tens of microns and is determined by requirement to have sufficient depth of focus (about 2-3 mm) This limitation is especially detrimental for endoscopic imaging of lumen structures that can have wide range of sizes and irregular lumen shapes. Methods of synthetic aperture radars (SAR) and sonars (SAS) imaging have been proposed for use in OCT to overcome limited depth of focus. Another example is given in U.S. Pat. No. 7,602,501 that discloses algorithms for three-dimensional inverse signal processing for full field optical coherent microscopy and for scanned-beam optical coherence microscopy (OCM). The use of teachings of U.S. Pat. No. 7,602,501 in endoscopic OCT imaging is impractical for several reasons. First, while phase stability of imaging in OCM can be recovered by re-processing signals using reflection from high-quality optical interface of a cover glass (as disclosed in U.S. Pat. No. 7,602,501), it is unpractical to fabricate high-quality reference surface in miniature endoscopic probes. Also image distortion caused by non-uniformity of rotation and especially pull-back makes algorithms of U.S. Pat. No. 7,602,501 not applicable for endoscopic OCT imaging. Second disadvantage of the method of U.S. Pat. No. 7,602,501 is that defocusing of optical energy in all directions results in significant loss of signal strength degrading image quality even if SAR signal processing is done with correct phases. Third, implementation of orthogonal scanning such as rotation and translation may be complicated for forward-looking imaging geometries in luminal structures. Therefore, it would be advantageous to apply methods of SAR signal processing to endoscopic imaging to mitigate limitations and trade-offs of Gaussian optics without above mentioned disadvantages of methods described in U.S. Pat. No. 7,602,501 as well as without need for fluoroscopic guidance and/or additional endoscopic guidance.
The Gaussian optics also imposes the trade-off between the spot size and the size of the probes. Namely, the smaller spot size and the larger working distance of a probe, the larger aperture and therefore the larger diameter of the probe should be. To overcome this limitation, as well as limitations of mechanical scanning, US2007/0188855 proposed methods of spectral encoding of spatial locations in tissue. The disadvantages of spectral encoding include decreased depth resolution of OCT imaging and increased complexity of the probes. Another imaging method, described in U.S. Pat. No. 7,474,407, was aimed at providing non-mechanical scanning and further miniaturization of probes and discloses an OCT apparatus having at least two fibers with adjustable phase delay between the fibers that claims advantages of non-mechanical scanning. However, U.S. Pat. No. 7,474,407 fails to disclose how exactly images are formed when phase delays are changed between the fibers. The method described in U.S. Pat. No. 7,474,407 B2 is based on use of TD-OCT and does not describe how it can be implemented with faster SD-OCT or FD-OCT that are more suitable for endoscopic imaging. The same patent document does not disclose any means to ensure stable interference between optical output from different fibers in endoscopic applications, when the states of polarization in different fibers will be arbitrary and sufficiently unstable due to temperature changes and twisting-and-bending of the probes. Finally, the use of two or more optical fibers prevents further miniaturization of probes distal ends. Thus it would be advantageous to provide probes that employ methods (of information encoding and/or non-mechanical scanning) that are without above mentioned limitations of US2007/0188855 A1 and U.S. Pat. No. 7,474,407 and that are devoid of fluoroscopic guidance and/or additional endoscopic guidance.
Another clinical aspect of using imaging probes in difficult-to-reach body lumens or body cavities is the need to deliver suction, irrigation, medication, or surgical tools to the region of interest during imaging procedures. The probes of related art appear to be not concerned with working channels to address this need. When used in working channel of endoscopes, imaging probes of related art have to be temporally removed in order to clear access to the region of interest via endoscopes working channel, thereby increasing the duration of the procedure and patient discomfort. In addition, the probes of related art have to be interchanged when imaging view has to be changed from side-view (associated with imaging of lumen's sub-surface wall tissue) to forward view (needed for image guidance and occlusion imaging), or when lumens of different sizes are being imaged. U.S. Pat. No. 7,706,646 discloses a multi-view probe head, the arrangement of which falls short of producing volumetric or easy-to-interpret cross-sectional images in the forward direction because the scanning pattern of this probe head is limited to conical regions in the forward direction. Additionally, the teachings of U.S. Pat. No. 7,706,646 rely on the use of polarization-maintaining fibers, which imposes a practical limit on miniaturization of the probes. Therefore what is needed is a probe or a set of probes with built in working channels (and method for the use of such probes) and means either to avoid changing probes or, at least, to facilitate such inevitable change.
Advantages of the OCT over ultrasound modalities are not limited to increased resolution in morphological structural images. The OCT also provides functional information of tissue physiology such as absorption, blood flow, and birefringence. The functional information about light absorption in tissue is important for concentration determination of various tissue chromophores. Light absorption can also be used advantageously to delineate lumen anatomy by exploiting differences in absorption characteristics between the lumen wall tissue and media filling the lumen. Examples of using the OCT to obtain functional information about light absorption in the tissue include the applications with oxygenated or de-oxygenated hemoglobin (as described, for example, in U.S. Pat. No. 6,015,969), or water content (as described in U.S. Pat. No. 6,134,003). However, U.S. Pat. No. 6,015,969 fails to describe whether its method can be applied to body lumens (as the method is based on TD-OCT and is not suitable for endoscopic applications), while U.S. Pat. No. 6,134,003 fails to disclose algorithms of determining the chromophores' concentrations from the OCT data. An imaging probe described in US 2011/0098572, which combines optical components with ultrasound transducers and cables with capability to map absorption features in tissue by using photoacoustic effects, is disadvantageous a far as lumen applications are concerned because of its increased complexity, as well as the size of a distal end that does not allow further miniaturization of the probe. Yet another disadvantage, already alluded to above, is a need for physical contact between probes and tissue and/or use of liquid-filled balloons as ultrasound imaging cannot be effective with air gaps between the probe and imaged tissue. Therefore it would be advantageous to have an imaging probe and an associated imaging console and algorithms capable of absorptive features imaging without deficiencies mentioned above as well as without need for fluoroscopic guidance and/or additional endoscopic guidance.